The present embodiments relate to generating an X-ray image of an object irradiated by X-ray radiation by an X-ray system including a counting digital X-ray detector.
X-ray systems are used for imaging for diagnostic examination purposes and for interventional procedures (e.g., in cardiology, radiology and surgery). Typically, X-ray systems have an X-ray tube and an X-ray detector, jointly mounted on a C-arm, for example, a high-voltage generator for generating the tube voltage, an imaging system, a system control unit, and a patient table. Biplane systems (e.g., C-arms) are likewise employed in interventional radiology. Generally, flat-panel X-ray detectors find application as X-ray detectors in many fields of medical X-ray diagnostics and intervention (e.g., in radiography, interventional radiology, cardioangiography, but also in therapeutic treatment applications for imaging within the context of monitoring and radiotherapy planning or mammography).
Flat-panel X-ray detectors in use today are generally integrating detectors and are based mainly on scintillators, composed of CsJ, for example, which convert X-ray radiation into comparatively low-energy radiation, visible light, for example. The light is converted into electrical charge in photodiode arrays. The photodiode arrays are then read out, typically row-by-row, via active control elements. FIG. 1 shows the basic structure of an indirectly converting flat-panel X-ray detector currently in use, having a scintillator 10, an active readout matrix 11 made of amorphous silicon or embodied in CMOS technology having a plurality of pixel elements 12 (e.g., with photodiode 13 and switching element 14) and drive and readout electronics 15 (see, for example, M. Spahn, “Flat detectors and their clinical applications”, Eur Radiol. (2005), 15: 1934-1947).
The active readout matrix 11 typically consists of a plurality of tiles or shingles (not shown). In order to generate large surface areas of, for example, 22×22 cm2, the matrix tiles are joined (e.g., ‘butted’) without gaps at the joints or else with a defined number of missing pixel rows or pixel columns. The function of the photodiodes 13 at the edge of such a tile is generally not constrained by their position. The scintillator 10 is generally embodied with a large surface area and covers a large number of tiles. This makes the overall behavior of the edge pixels in present-day integrating flat-panel detectors well-mannered such that either few or no corrections at all are to be made at the joints of the matrix tiles. Thus, in the event of defined pixel rows being missing between two matrix tiles, for example, trivial interpolation algorithms are employed in order to reconstruct the missing image information (e.g., the defect correction method described in the publication U.S. Pat. No. 6,763,084 B2) when, as described, defects occurring dynamically and not primarily at a fixed location are also handled.
Depending on beam quality, the quantum efficiency for a CsJ-based scintillator having a layer thickness of, for example, 600 μm ranges between approximately 50% and 80% (see, for example, M. Spahn, “Flat detectors and their clinical applications”, Eur Radiol (2005), 15: 1934-1947). The spatial frequency dependent DQE(f) (Detective Quantum Efficiency) is upwardly limited as a result and lies significantly thereunder for typical pixel sizes of, for example, 150 to 200 μm and for the spatial frequencies of 1 to 2 lp/mm (e.g., line pairs per mm) that are of interest for the applications. In order to enable new applications (e.g., dual-energy, material separation, etc.), but also to increase the quantum efficiency further, use is increasingly being made of the potential of counting detectors or energy-discriminating counting detectors mainly based on direct-converting materials (e.g., such as CdTe or CdZTe=CZT) and contacted Application-Specific Integrated Circuits (ASICs) (e.g., implementation in CMOS technology). Other materials such as Si or GaAs may likewise be of interest for certain applications. Counting detectors count the incident X-ray quanta individually, instead of integrating the X-ray quanta as a whole, with the result that electronic noise may be suppressed almost completely.
An example layout of such counting X-ray detectors is illustrated in FIG. 2. X-ray radiation is converted in the direct converter 224 (e.g., CdTe or CZT), and the generated charge carrier pairs are separated via an electrical field that is generated by a common top electrode 26 and a pixel electrode 25. The charge generates a charge pulse in one of the pixel electrodes 25 of the ASIC 27 that are implemented as pixel-shaped. The height of the charge pulse corresponds to the energy of the X-ray quantum, and the charge pulse is registered as a count event if the charge pulse exceeds a defined threshold value. The threshold value serves to differentiate an actual event from electronic noise or, for example, also to suppress k-fluorescence photons in order to avoid multiple counts. The ASIC 27, a corresponding section of the direct converter 224, and a coupling between direct converter 224 and ASIC 27 (e.g., by bump bonds 36 in the case of direct-converting detectors) in each case form a detector module 35 having a plurality of pixel elements 12. The ASIC 27 is arranged on a substrate 37 and is connected to peripheral electronics 38. A detector module 35 may also have one or more ASICs 27 and one or more part-pieces of a direct converter 224, chosen according to requirements in each case.
Many of the direct converters 224 that promise high signals and count rates, such as CdTe or CZT, may only be fabricated at reasonable cost in small surface areas (e.g., 2×2 cm2 or 3×3 cm2). ASICs 27 having a complex pixel structure, such as are required for counting detectors, may likewise only be produced with an acceptable yield in small surface areas. By greater investment of resources, somewhat larger surface areas such as, for example, 2×8 cm2 or 3×6 cm2 may be provided, such that, for example, four 2×2 cm2 or two 3×3 cm2-sized direct converters 224 may be mounted onto the corresponding ASICs 27 in order to form a detector module 35 in combination. In any case, such detector modules 35 are nevertheless small compared to the overall size of an average flat-panel image detector, such as is required for applications in angiography (e.g., 20×20 cm2 or 30×40 cm2). In order to obtain a sufficiently large counting X-ray detector, a plurality of detector modules 35 are aligned next to one another or in a matrix-like arrangement (e.g., on four sides in the case of rectangular/square-shaped detector modules).
The general schematic layout of a counting pixel element 12 is shown in FIG. 4. The electrical charge is collected in the pixel element 12 via the charge input 28 and amplified in the pixel element 12 with the aid of a charge amplifier 29 and a feedback capacitor 40. In addition, the pulse shape may be adjusted at the output in a pulse shaper (e.g., filter) (not shown). An event is counted by incrementing a digital memory unit 33 (e.g., adder or counter) by one when the output signal exceeds a selectable threshold value. This is detected via a discriminator 31. The threshold value may also be specified as a fixed analog value, but is generally applied via a digital-to-analog converter (DAC) 32 and is consequently variably adjustable over a certain range. The threshold value may be selectable either locally pixel-by-pixel, via the discriminator 31 (e.g., local discriminator) and the DAC 32 (e.g., local DAC), as shown, or globally for a plurality of/all pixel elements via a global discriminator and DAC, for example. Counts may subsequently be read out via a drive and readout unit or via peripheral electronics 38.
Over and above a global DAC, which serves, for example, for setting a specific keV threshold for an entire detector module or the entire X-ray detector, a further pixel-by-pixel matching may be provided in order to correct pixel-to-pixel fluctuations (e.g., fluctuations in amplifiers 29, local material inhomogeneities of the detector material, etc.). This pixel-by-pixel calibration or correction DAC generally has a much higher resolution than the global DAC and is settable, for example, over a keV range within which the pixel-to-pixel fluctuations are expected (e.g., 6 keV). If such a calibration or correction DAC is provided, then it is advantageous to perform the global DAC and the correction DAC separately on account of the cited different resolutions. The global DAC may then be configured with rather a lower resolution (e.g., 2 keV/bit), which generates a voltage that is present at each pixel element of the detector module or all detector modules of a detector and onto which a pixel-by-pixel correction voltage is superimposed pixel by pixel by way of a higher-resolution correction DAC (e.g., 0.1 keV/bit or 0.5 keV/bit). If a plurality of threshold values and counters are provided per pixel element (e.g., spectral imaging), then a plurality of global DACs are provided. A calibration or correction DAC may be provided for each discriminator if, for example, the circuit exhibits a non-linear response.
FIG. 5 shows a schematic for an entire array of counting pixel elements 12 (e.g., 100×100 pixel elements of 180 μm each). In this example, it would have a size of 1.8×1.8 cm2. For large-area X-ray detectors (e.g., 20×30 cm2), a plurality of detector modules 35 are combined with one another (e.g., approximately 11×17 modules would produce this surface area) and are connected via the common peripheral electronics. For the connection between ASIC 27 and peripheral electronics 38, use is made of Through Silicon Via (TSV) technology, for example, in order to provide the detector modules 35 are aligned next to one another in the tightest possible four-sided arrangement.
In the case of counting and energy-discriminating X-ray detectors, two or more different threshold values per pixel (e.g., four), as shown in FIG. 6, are introduced by four pairs made up of DAC 32 and discriminator 31, and the height of the charge pulse, corresponding to the predefined threshold values (e.g., discriminator threshold values), is classified into one or more of the digital memory units 33 (counters). The X-ray quanta counted in a specific energy range may then be obtained by calculation of the difference between the counter contents of two corresponding counters. The discriminators 31 may be set, for example, with the aid of digital-to-analog converters (DAC) 32, for the entire detector module or pixel by pixel within given limits or ranges. The counter contents of counting pixel elements 12 are read out module by module in succession via a corresponding readout unit.
Various effects may now lead to a situation where an absorbed X-ray quantum deposits energy not just in one pixel, but where a portion of the energy is deposited in the neighboring pixels due to processes such as charge sharing or fluorescence photons (e.g., k-fluorescence). This may lead to miscounts (e.g., to multiple counts or no count) if the respective deposited energies lie below the threshold values set at the pixels or also to incorrect assignment of the energy in the case of energy-discriminating detectors. In order to solve the above-described problem, summation and anticoincidence circuits in which the charge deposited in neighboring pixels within a given time interval (e.g., coincidence) is added together and assigned to a specific pixel and the summation signal is compared with the threshold value of a discriminator or a plurality of discriminators of the pixel may be used.
One of the critical points in the case of counting detectors is the modular structure. As already mentioned, the detector modules of a counting detector are significantly smaller than the detector modules of integrating detectors. In order to obtain a sufficiently large X-ray detector, a plurality of detector modules are aligned next to one another or in a matrix-like arrangement (e.g., on four sides in the case of rectangular/square-shaped detector modules).
For mechanical and thermal reasons (e.g., precision with which the modules may be produced, dimensions, etc.), a gap, for example, in the form of rows or columns of pixels (e.g., in the x- and/or y-direction) that are insensitive to X-ray radiation, generally arises unintentionally between adjacently arranged modules. This may lead, for example, to one or more rows of missing pixels in each case between the modules. In a design that manages without missing pixels, the edge pixels of detector modules exhibit a different signal response due to the design compared to pixels arranged centrally on a detector module. This is attributable, for example, to the deformation of the electrical field at the edge of the X-ray converter due to the missing neighboring pixels on at least one side (known as discontinuity). For example, a lower efficiency of the charge collection in the edge pixels may lead to lower pulse heights, and this may lead to a changed count rate behavior above the defined energy thresholds. Given otherwise comparable conditions (e.g., with homogeneous X-ray flux and the same location-independent X-ray spectrum), the noise performance at the edge of a detector module may be changed by comparison with more centrally arranged pixels. In the case of ASICs that support energy summation and anticoincidence circuits of neighboring pixels, there is the problem that this generally may not be realized beyond ASIC boundaries. In other words, pixels arranged at peripheral positions on a detector module may also behave differently from centrally arranged pixels in terms of energy deposition and energy restoration. Further, fluorescence photons may escape more easily at the edges of the direct converter because no reabsorption may take place there due to the absence of material. Depending on X-ray beam quality and/or object to be examined, scatter radiation effects may manifest themselves differently at the module edges than in the center of the module. Also, given a comparable incident X-ray flux, an effectively smaller active surface of edge pixels of a detector module by design may generally lead to a smaller count rate and lower signal-to-noise ratio. The active surface area of edge pixels may be reduced in size due to a guard ring, for example. A guard ring in the form of an electrode defines the properties at the detector module edges and in certain cases improves the properties of peripherally arranged pixels.
A further problem may arise because electrical contacts are to be brought out vertically rather than laterally through the silicon of the ASIC chip with the aid of TSV technology, thereby providing the electrical connection (e.g., voltage supply of the chips, control and data lines) to the underlying electronics. Such TSVs use vertical openings that, depending on silicon thickness, are, for example, 100 μm to 200 μm in size and accordingly may assume pixel size. Surface area within the detector is likewise not available for the signal generation.
The quality of the resulting X-ray images suffers markedly due to the increased number of discontinuities occurring as a result of the small-scale granularity of the modular structure of counting detectors.